A special requirement for MRI is a strong uniform magnetic field, typically 0.2 to 2 Tesla, with a field homogeneity of a few parts per million in the imaging volume, typically a sphere of 30 cm to 50 cm in diameter. Most commonly such a field is produced by an electromagnet having a solenoid construction but this necessitates a patient being surrounded by the magnet and enclosed within a central bore tube. Although the length of this bore tube is typically 1.6 m or less it can cause a feeling of claustrophobia which is extremely distressing to some patients. Furthermore, such a design does not give any access to the patient, which may be essential for interventional procedures, or when the proximity of a companion is required to put the patient at ease. However by using open magnets, these problems are overcome, or at least substantially mitigated.
Open electromagnets for use in MRI systems are well known. One form of an open electromagnet, often described as a ‘split pair’ comprises a pair of juxtaposed sets of coils, which are generally of a solenoid construction and may include a bore tube around the axis. The sets of coils are held apart by a support structure with a gap between the assemblies wide enough for a person to stand, so that access may be gained to the imaging volume between them along any of the principal axes of the system. Normally, the patient would be positioned along the axis of the solenoids. The gap between the solenoid assemblies is adequate for access to do interventional work. Such a system is described in U.S. Pat. No. 5,381,122 and in a paper by Laskaris et al, entitled ‘A Cryogen-Free Open Superconducting Magnet for Interventional MRI Applications’ published in IEEE Transactions in Applied Superconductivity, Volume 5, No. 2, June 1995.
Another solution is a design wherein two sets of solenoidal coils are concentric with a vertical axis. Such a design is described in, for instance, U.S. Pat. No. 5,874,882. However, the disadvantage of such a design is that for a given central field strength, a large stray field is produced. The size of the stray field can be reduced by increasing the size of the so-called shielding coil. However, an increase of the size of the shielding coils will reduce the central field. This reduction, in turn, requires a substantial increase in the size of the driving coils. The result is that a magnet, comprising concentric axial coils, spaced apart to allow good access to the patient, will require a substantial amount of conductor for a given central field to keep the stray field to a reasonable size.
In Huson et al, PCT WO93/15514, an electromagnet is disclosed which comprises a pair of juxtaposed magnetic poles of opposite polarity between which the imaging volume is defined, which poles are linked and supported by an iron yoke or superconducting solenoid which provide a magnetic flux return path. The iron yoke principally comprises a generally C-shaped iron frame. Because large amounts of iron are required, these known C-shaped magnets are very heavy, especially for high field magnets which require many tonnes of iron to define the flux return path. Huson et al further discloses the use of shielding coils for constraining the magnetic return flux but the amount of conductor used, and hence the cost, is high.
In each of these known open magnets the direction of the magnetic field is along the axis of the coils, and therefore perpendicular to the plane of the gap. The magnetic force between the juxtaposed magnetic poles of these known open magnets is very large, and acts in a direction to close the gap. This imposes large compressive forces on a structure used to support the poles, and requires that the structure has adequate strength and stability to resist the forces. The structure must therefore be substantial, and provides a significant impediment to free access to the imaging volume,